The present invention relates to medical imaging systems. It finds particular application in the calibration of positron emission tomography (PET) or coincidence-capable gamma camera medical imaging systems, and will be described with particular reference thereto. It is to be appreciated that the invention may also find application in other types of nuclear cameras as well as other types of diagnostic imaging with two dimensional rotatable detectors.
A number of types of radiological events produce two emission photons which travel outward in exactly opposite directions. For example, a positron-electron annihilation event produces such an oppositely directed photon pair. Coincidence imaging systems take advantage of this geometric property in spatially localizing the radiological event to a line of response (LOR). The LOR is defined as the line connecting two simultaneous radiation detection events, which are presumed to correspond to detection of the two oppositely directed photons. Two detection events are typically judged to be simultaneous, or coincident, if both detections occur within a preselected coincidence time window. The radiological event, e.g. the positron-electron annihilation, will have occurred at some point along the LOR. Ideally, for an approximate point source which generates many positron-electron annihilation events within a very confined space, the LOR corresponding to each annihilation event will pass through the point source, and the point source may thus be located in two-dimensional or three-dimensional space, dependent upon the geometry of the detector system. Similarly, an extended radiological source, such as an organ which has absorbed an appropriate administered radiopharmeceutical, may be studied by reconstructing the LOR""s corresponding to electron-positron annihilation events generated by the radiopharmeceutical into an image.
Non-idealities in real detection systems, such as mechanical misalignments, degrade image resolution. For example, misalignment of the radiation detectors blurs the image. The precision with which the radiation detectors can be mechanically aligned is limited by the weight and bulkiness of the shielded detectors, and the alignment precision required for adequate image resolution is often not practically attainable by purely mechanical procedures. Alignment-related resolution problems are particularly acute for positron coincidence imaging using a multi-head single photon emission computed tomography (SPECT) gamma camera with coincidence circuitry, often referred to as a gamma-PET imaging system. The gamma-PET provides a versatile and less expensive alternative to a dedicated PET imaging system, but the typically large heads and rotating gantry complicate mechanical alignment.
Improved resolution can be obtained by following the mechanical alignment with a calibration step. The calibration step preferably determines correction factors for the positional coordinates of each detector. The correction factors may subsequently be applied during data acquisition or analysis to improve image resolution. Gamma-PET systems have been calibrated using a SPECT data acquisition mode, the results of which are applied when operating the system in a PET or SPECT imaging mode. This approach corrects for the tangential detector head coordinate. However, SPECT is not strongly sensitive to the radial and orientational detector coordinate parameters. An additional disadvantage of calibration in SPECT mode is that subsequent to calibration the detector heads are re-positioned into the PET configuration, which may introduce additional misalignment not accounted for by the prior SPECT calibration.
The present invention contemplates a calibration procedure for coincidence imaging systems such as gamma-PET and dedicated PET systems, which overcomes the above shortcomings and others.
In accordance with one aspect of the present invention, a method for calibrating a coincidence imaging system which includes a plurality of radiation detectors is disclosed. A plurality of coincidence radiation events associated with a point radiation source are measured. Initial values are assigned for a set of fitting parameters. A minimization algorithm is applied, which includes calculating lines of response (LOR) based upon the fitting parameters and the measured radiation events, generating a figure of merit characterizing the apparent size of the point radiation source based upon the LOR""s, and optimizing the fitting parameters to produce a minimized figure of merit. After the minimization, a correction factor relating to a positional coordinate of a detector is extracted from the fitting parameters.
In accordance with another aspect of the present invention, a method for imaging using a plurality of radiation detectors is disclosed. A plurality of coincidence radiation events associated with a point radiation source are measured. Initial values are assigned for at least one fitting parameter. Lines of response (LOR) are calculated based upon the at least one fitting parameter and the measured radiation events. A figure of merit is generated that characterizes the apparent size of the point radiation source based upon the LOR""s. The at least one fitting parameter is optimized using a minimization algorithm which includes iteratively repeating the calculating and generating steps to produce a minimized figure of merit. At least one correction factor is extracted from the at least one optimized fitting parameter. A set of radiation data is acquired from an associated subject. The radiation data is corrected for camera misalignment by correcting the spatial coordinates of the detected radiation events using the at least one correction factor. An image representation is reconstructed from the corrected radiation data.
Preferably, the at least one fitting parameter includes a parameter related to the radial positional coordinate of a detector, a parameter related to the tangential positional coordinate of a detector, and a parameter related to the orientational positional coordinate of a detector. For a multiple-head imaging system, the fitting parameters preferably include: xcex94ri, i=1 to N, where xcex94ri is a correction for the radial coordinate of the ith detector; xcex94tj, j=1 to N, where xcex94tj is a correction for the tangential coordinate of the jth detector; and xcex94xcex8k, k=2 to N, where xcex94xcex8k is a correction for the orientational coordinate of the kth detector.
The figure of merit is preferably generated by summing the distance or the square of the distance of closest approach of each LOR to a spatial point, in which case the positional coordinates of the spatial point are fitting parameters. Preferably, LOR""s whose distance of closest approach is greater than a preselected distance are discarded. Alternatively, the figure of merit is generated by obtaining the crossing point or the distance of closest approach of each pair of LOR""s and calculating the standard deviation of the crossing points or the obtained distances.
In accordance with yet another aspect of the present invention, a method of PET imaging is disclosed. Coincidence radiation events from a calibration source are detected with at least two detector heads. Correction factors that correct for mechanical misalignment of the detector heads are calculated from the coincidence detected calibration source radiation. During a diagnostic imaging procedure performed on a subject, image data are generated in response to radiation collected with the detector heads. The image data are corrected with the correction factors. The corrected image data are reconstructed into an image representation.
In accordance with still yet another aspect of the present invention, a coincidence imaging system is disclosed. The system includes a gantry. A plurality of flat panel detectors are disposed about the gantry. A data memory stores measured data about radiation events detected by the detectors. A calibration memory stores a plurality of calibration parameters for correcting data measured during a patient scan. A processor in communication with the calibration memory and with the data memory calculates the calibration parameters by a minimization algorithm that includes optimizing fitting parameters with respect to acquired radiation data associated with a point radiation source.
Preferably, the calibration parameters include parameters relating to positional coordinates of the plurality of detectors. The gantry is preferably rotatable. The figure of merit is preferably generated by summing the square of the distance of closest approach of each LOR to a spatial point, in which case the positional coordinates of the spatial point are fitting parameters. Alternatively, the figure of merit is generated by obtaining the crossing point of each pair of LOR""s and calculating the variance of the crossing points. To reduce noise, the minimization algorithm preferably discards measured data about radiation events whose energy is outside a preselected energy range.
One advantage of the present invention is that the calibration is performed in coincidence imaging mode and does not require mechanical re-configuration the system between calibration and PET imaging.
Another advantage of the present invention is that it simultaneously calibrates all three detector head positional coordinates, e.g. the radial, tangential, and orientational coordinates.
Another advantage of the present invention is that it provides improved image quality.
Yet another advantage of the present invention is that it is compatible with both gamma-PET and dedicated PET systems.
Still further advantages and benefits of the present invention will become apparent to those of ordinary skill in the art upon reading and understanding the following detailed description.